An enormous expenditure of health-care resources was required for the repair and replacement of diseased tissue structures and organs. The most common treatment, replacement with an autograft, produces less than optimal results. However, the supply of autograft, and even allograft, is very limited. Engineering tissues and organs with mammalian cells and a scaffolding material as emerged as a promising alternative approach in the treatment of malfunctioning or lost organs compared to the use of harvested tissues and organs (see Langer, R. S. and J. P. Vacanti, “Tissue engineering: the challenges ahead,” Scientific American 280(4), 86 (1999)). In this approach, a temporary scaffold is needed to serve as an adhesive substrate for the implanted cells and a physical support to guide the formation of the new organs. Accordingly, the scaffold materials must be custom-engineered to match the biomechanical, biochemical, and biological needs of the specific tissue or organ they are designed to replace. Transplanted cells adhere to the scaffold, proliferate, secrete their own extracellular matrices (ECM), and stimulate new tissue formation (see Langer, R. and J. Vacanti, “Tissue Engineering”, Science 260 (5110), 920-926 (1993); Hubbell, J. A., “Biomaterials in Tissue Engineering”, Bio/Technology 13, 565 (1995); and Saltzman, W. M., “Cell interactions with polymers,” Principles of Tissue Engineering, R. Lanza, R. Langer, and W. Chick, Editors (1997) Academic Press, R. G. Landes Company, Austin, Tex., 225). During this process, the scaffold must be degraded and eliminated gradually and disappears eventually. Therefore, in addition to facilitating cell adhesion, promoting cell growth, and allowing the retention of differentiated cell functions, the scaffold should be biocompatible, biodegradable, highly porous with a large surface/volume ratio, mechanically strong, and malleable into desired shapes. The biophysical and biomechanical properties of a biomaterial scaffold are crucial for the outcome of tissue engineering. In many circumstances, the material selection is a compromise among the many physical and biological requirements. Synthetic biodegradable polymers have been attractive candidates for scaffolding materials because they degrade as the new tissues are formed, eventually leaving nothing foreign to the body. And their physical and biological properties can be controlled and tailored through different synthetic conditions and methods.
Aliphatic polyesters are one class that consists of synthetic biodegradable polymers, such as poly(glycolic acid) (PGA), poly(lactic acid) (PLA), and their copolymer of poly-(DL-lactic-co-glycolic acid)(PLGA). PLA, PGA, and PLGA have been approved by the U.S. Food and Drug Administration for some human clinical applications, such as surgical sutures and implantable devices. One of their potential advantages is that their degradation rate can be adjusted to match the rate of regeneration of the new tissue. With sufficient mechanical strength, they can keep this framework until the new tissue forms. They also can be fabricated to be the same complicated shapes or structures as the tissues or organs to be replaced.
Although such synthetic materials are widely used, they still have some disadvantages, such as hydrophobicity, the lack of cell-recognition signals, etc. These results show that there is no sufficient cell adhesion on the surface of these polymer materials. Their interactions with the host environment still have much potential for improvement. How to improve the biomaterial/cell interaction for eliciting the controlled cellular adhesion and maintaining differentiated phenotypic expression has become one of the major challenges in the field of tissue engineering.
To overcome the drawbacks associated with synthetic biodegradable polymer materials, much attention has focused on coating cell adhesion-enhancing agents on these materials, such as collagen, bone-like apatite, hydroxyapatite, on the surface of polymer materials.
It is an object of the present invention to overcome the disadvantages and problems in the prior art.
The following description of certain exemplary embodiment(s) is merely exemplary in nature and is in no way intended to limit the invention, its application, or uses.
Now, to FIGS. 1-5,
FIG. 1 is an embodiment of the method of the present invention. As a first step, keratin is mixed with a biomaterial 101. Keratins are the major structural fibrous proteins constructing hair, wool, nail, and so on, which are characteristically abundant in cysteine residues (7-20% of the total amino acid residues). As alternative natural proteinous biomaterials for collagen, wool keratins have been demonstrated to be useful for fibroblasts and osteoblasts, owing to their cell adhesion sequences, arginine-glycine-aspartic acid (RGD) and leucine-aspartic acid-valine (LDV), biocompatibility for modification targets. Moreover, they are biodegradable in vitro (by trypsin) and in-vitro (by subcutaneous embedding in mice). FIG. 1 is an embodiment of a scanning electron microscope image of wool keratin particles.
The biomaterial can be natural or synthetic, and is selected from the group consisting of poly(lactide-co-glycolide)(PLGA), poly(lactide)(PLLA), polyglycolic acid (PGA), polyanhydrides, poly(ortho ethers), poly caprolactone, polyethylene glycol (PEG), polyurethane, copolymers thereof, and mixtures thereof.
The keratin-biomaterial solution is then mixed with porogen 103. Porogen is selected from the group consisting of sodium chloride, sodium sulfate, potassium chloride, sodium iodide, sodium fluoride, potassium fluoride, sodium nitrate, sodium iodate, mixtures thereof, sodium hydroxide, fructose, saccharin, glucose, mixtures thereof, paraffin, beeswax, mixtures thereof, naphthalene, and gelatins. The porogen may be formed into any shape as desired and/or necessary. In a preferred embodiment, the predetermined shape is selected from the group consisting of cubic or other geometrically shaped crystals, spheres, fibers, discs, regular geometric shapes, irregular geometric shapes, and mixtures thereof.
The solution and porogen are then dried 105, and then the porogen is dissolved or leached out from the solution 107.
It is to be understood that the cell adhesion-enhancing keratin used in the present invention may include a liquid. Preferably, the cell adhesion-enhancing keratin when mixed is at least one of a solution, a suspension, a melt, a slurry, flowable powders, flowable fibers, flowable pastes, and mixtures thereof. It is to be understood that the natural or synthetic biomaterials may be any composition which flows adequately for blending purposes. In one preferred embodiment, the liquid is a solvent and the biomaterial, natural or synthetic, is a polymeric composition.
The method of the present invention may be performed continuously, i.e., it may be automated wherein the porogen is used to print the 3-D structure, or in batches, manually or automatically.